Nuclear magnetic resonance (NMR) is associated with coherent resonant rotation of magnetic atomic nuclei, around an axis parallel to a large static applied magnetic field, following excitation by a radio-frequency pulse. The rotation frequency (the Larmor frequency) is proportional to the applied magnetic field, given by f=γB, where γ=42 MHz/T for the hydrogen nucleus (proton) and of similar magnitude for other nuclei. The coherent decay time, known as T2, is typically of order 100 ms, reflecting the very high-Q nuclear rotations with only weak interactions with the atomic environments. In magnetic resonance imaging (MRI), appropriate weak gradients in applied magnetic field are used to obtain frequency gradients, thus enabling Fourier transforms of the nuclear signal to be mapped onto spatial images. However, the coherent signal is extremely weak, corresponding to radio frequency (RF) magnetic fields (for 1-mm resolution) of order 10−14 T (=10 fT) in an applied field of order 1 T. For an applied field in the z-direction, the RF magnetic signal rotates in the x-y plane at the Larmor frequency, corresponding to circular polarization (see FIG. 1).
Magnetic resonance imaging (MRI) is primarily a medical imaging technique commonly used in radiology to visualize detailed internal structures and limited functions of the body, but also used in non-destructive testing and other fields. MRI is useful because it provides great contrast between the different soft tissues of the body, such as lipid and aqueous. MRI is discussed in detail at en.wikipedia.org/wiki/Magnetic_resonance_imaging (incorporated herein by reference).
In addition to the desired RF signal, the measured RF fields also comprise noise within the same frequency band. There may be many sources of noise, including noise associated with electrical currents in the antenna and amplifiers in the receiver. The dominant source of noise in MRI systems typically comprises fluctuating Johnson noise associated with thermally excited eddy currents in the electrically conducting medium of the human body or other object under examination. Johnson noise is the electrical noise generated by the thermal agitation of charge carriers, such as electrons and ions, inside an electrical conductor at equilibrium. Johnson noise is approximately white, meaning that its spectral density is nearly equal throughout the frequency spectrum, and therefore including even the relatively narrow bands of interest during MRI signal processing. See, for example, “Signal-to-noise ratio in MRI”, T. W. Redpath, British Journal of Radiology, vol. 71, pp. 704-707 (1998), incorporated herein by reference. This Johnson noise is largely independent of the static magnetic field, while the signal of interest is proportional to the applied field. (This is the main impetus for the use of very large magnetic fields.) These eddy currents create broadband RF magnetic fields which also couple to the pickup coil or coils, with a form within a narrow band of interest that may be expressed as Bn(t)=Bn0 cos(ωt+ϕn), where Bn0 and ϕn are the amplitude and phase of the fluctuating noise component of the magnetic field that couples to the coil. MRI receivers generally filter out all noise outside of the narrow band required for the MRI signal. For example, a typical receiver bandwidth might be 50 kHz or less. In this case, Bn0 and ϕn can vary significantly over a time greater than about 20 microseconds limited by the receiver bandwidth. Note that the nuclear signal itself is coherent for a time T2 of order 100 ms, while the various noise sources are incoherent (or coherent over much shorter timescales), so that signal integration generally increases the signal amplitude linearly with the time, while it increases the noise amplitude with the square root of the time. The signal-to-noise ratio (SNR) thus increases with averaging or measurement repetition. However, performing an integration over the coherence time to maximize the signal to noise ratio may unduly prolong an MRI scan, which is uncomfortable to the patient, and may lead to movement artifacts during the scan.
There has been relatively little effort in the prior art devoted to measuring and characterizing the body noise in MRI. However, U.S. Pat. No. 6,865,494, expressly incorporated herein by reference, proposes to provide “noise tomography”, in which a three dimensional scan of Johnson noise is itself the output, wherein tissues having different conductivity have variations in the measured noise.
One type of MRI antenna is described in by Eydelman in U.S. Pat. No. 6,636,040, incorporated herein by reference. Eydelman's antenna, and that of similar devices, reads both the signal, as well as noise from Eddy Currents caused by various dynamic magnetic fields in the body, including those induced by the MRI machine itself. The isolation of the signal with respect to in-band noise is difficult and often impossible, leading to impaired resolution or aberrations, longer scan time, and/or image noise.
Consider a planar pickup coil in the y-z plane, which detects a field component of the rotating nuclear magnetic signal Bsx(t)=Bs0 cos(ωt) in the x-direction, where w=2πf is the Larmor frequency. An identical coil oriented in the x-z plane will detect the corresponding nuclear magnetic field component Bsy(t)=Bs0 sin(ωt) in the y-direction, shifted by 90° from the first coil. Taken together, the two coils form a quadrature antenna, which is known in the prior art. The prior art further teaches that applying a 90° phase shift and adding the signals from the two antennas will increase the nuclear magnetic signal amplitude by a factor of two. See, for example, U.S. Pat. Nos. 7,649,353; 4,769,605; 5,351,688.
If one has two antennas in a quadrature configuration, adding the two signals with a 90° phase shift increases the signal amplitude by a factor of two, as noted above. If the noise signals Bn from the two antennas are uncorrelated, then adding the two signals increases the noise amplitude by a factor of the square root of two (√2). So the SNR would increase by √2, or about 3 dB in terms of power ratios. Such an improvement is significant but limited.
A further aspect of the prior art is the development of more sensitive low-noise receivers, including cryogenic coils, superconducting sensors (based on superconducting quantum interference devices or SQUIDs), and low-noise amplifiers. These may be useful when the signal of interest is especially weak, as for example in relatively low magnetic fields. See, for example, U.S. patent application Ser. No. 12/954,291, filed Nov. 24, 2010, expressly incorporated herein by reference in its entirety. See also, for example, U.S. Pat. Nos. 6,885,192; 7,053,610; 6,538,445; 5,276,398; see also L. Darasse and J. C. Ginefri, “Perspectives with cryogenic RF probes in biomedical MRI”, Biochimie, vol. 85, p. 915 (2003) and “Calculated SNR of MRI detected with SQUIDs and Faraday detectors”, W. Myers, et al., Journal of Magnetic Resonance, vol. 186, p. 182 (2007), incorporated herein by reference. However, in many cases the receiver noise is already less than the body noise, so that very little additional SNR is obtained from further reduction in receiver noise.
Superconducting quantum interference devices (SQUIDs) are very sensitive magnetometers used to measure extremely weak magnetic fields, such as those produced during MRI medical tests, based on superconducting loops containing Josephson junctions.
See, U.S. Pat. Nos. 7,688,069, 7,671,587, 7,603,158, 7,573,268, 7,573,264, 7,560,289, 7,535,228, 7,525,314, 5,586,064, 3,801,877, 7,521,928, 7,482,807, 7,474,095, 7,466,132, 7,395,107, 7,363,070, 7,248,044, 7,218,104, 7,197,352, 7,193,415, 7,187,169, 7,144,376, 7,130,675, 7,123,952, 7,092,748, 7,038,450, 6,897,654, 6,865,494, 6,697,660, 6,681,131, 6,544,170, 6,522,908, 6,477,398, 6,374,131, 6,370,414, 6,208,884, 6,187,032, 6,159,444, 6,150,809, 6,073,040, 6,031,373, 6,002,254, 5,982,174, 5,827,501, 5,771,894, 5,771,893, 5,755,227, 5,752,514, 5,682,889, 5,671,740, 5,657,756, 5,608,320, 5,601,081, 5,600,243, 5,594,849, 5,543,770, 5,495,849, 5,442,290, 5,426,365, 5,408,178, 5,384,109, 5,351,006, 5,339,811, 5,326,986, 5,325,854, 5,303,705, 5,274,331, 5,233,992, 5,208,534, 5,187,327, 5,057,776, 5,021,739, 4,951,674, 7,116,102, 7,053,610, 6,885,192, 6,724,188, 4,390,840, 4,442,404, 4,573,015, 4,588,947, 4,851,777, 4,864,237, 4,906,931, 4,987,368, 5,057,776, 5,208,533, 5,254,950, 5,300,887, 5,343,147, 5,557,199, 5,600,243, 5,835,995, 6,023,161, 6,031,373, 6,159,444, 6,544,170, 6,724,188, 4,879,516, 4,695,801, 7,126,333, 6,838,875, 5,436,564 and 2006/0186882, each of which is expressly incorporated herein by reference.
The prior art has not effectively solved the issue of noise in bioelectric and/or biomagnetic field measurements, such as the background body noise in MRI systems, so that one may obtain high-resolution images with ultra-sensitive receivers, without requiring the largest magnetic fields and long integration times.